Biosensor device and method

ABSTRACT

A biosensor apparatus for detecting a binding event between a ligand and receptor. The apparatus includes an electrode substrate coated with a high-dielectric hydrocarbon-chain monolayer, and having ligands attached to the exposed monolayer surface. Binding of a receptor to the monolayer-bound ligand, and the resultant perturbation of the monolayer structure, causes ion-mediated electron flow across the monolayer. In one embodiment, the monolayers have a coil-coil heterodimer embedded therein, one subunit of which is attached to the substrate, and the second of which carries the ligand at the monolayer surface.

This application claims the priority of U.S. Provisional application No.60/016,196 filed Apr. 25, 1997, which is incorporated herein byreference.

FIELD OF THE INVENTION

The present invention relates to biosensors, in particular, to abiosensor for measuring a binding event between a ligand and aligand-binding receptor, and to methods employing such biosensor.

BACKGROUND OF THE INVENTION

To a great extent, diagnostic tools used for detecting or quantitatingbiological analytes are based on ligand-specific binding between aligand and a receptor. Ligand-receptor binding pairs used commonly indiagnostics include antigen-antibody, hormone-receptor, drug-receptor,cell surface antigen-lectin, biotin-avidin, and complementary nucleicacid strands, wherein said ligand is typically the smaller of the twobinding pair members. The analyte to be detected may be either member ofthe binding pair; alternatively, the analyte may be a ligand analog thatcompetes with the ligand for binding to the complement receptor.

A variety of methods for detecting ligand/receptor interactions havebeen developed. The simplest of these is a solid-phase format employinga reporter-labeled ligand whose binding to or release from a solidsurface is triggered by the presence of analyte ligand or receptor. In atypical solid-phase sandwich type assay, for example, the analyte to bemeasured is a ligand with two or more binding sites, allowing ligandbinding both to a receptor, e.g., antibody, carried on a solid surface,and to a reporter-labeled second receptor. The presence of analyte isdetected (or quantitated) by the presence (or amount) of reporter boundto solid surface.

In a typical solid-phase competitive binding assay, an analyte ligand(or receptor) competes with a reporter-labeled analyte analog forbinding to a receptor (or ligand) carried on a solid support. The amountof reporter signal associated with the solid support is inverselyproportional to the amount of sample analyte to be detected ordetermined.

The reporter label used in both solid-phase formats is typically avisibly detectable particle or an enzyme capable of converting asubstrate to an easily detectable product. Simple spectrophotometricdevices allow for the quantitation of the amount of reporter label, forquantifying amount of analyte.

Detecting or quantitating ligand-specific binding events is alsoimportant in high-throughput methods being developed for combinatoriallibrary screening. In a typical method, a large library of possibleeffector molecules (ligands) is synthesized. The library members arethen screened for effector activity by their ability to bind to aselected receptor. The approach has the potential to identify, forexample, new oligopeptide antigens capable of high-specificity bindingto disease related antibodies, or small-molecule compounds capable ofinteracting with a selected pharmacological target, such as a membranebound receptor or cellular enzyme.

High-throughput screening methods typically employ simple liganddisplacement assays to detect and quantitate ligand binding to areceptor. Displacement assays have the advantage of high sensitivity,e.g., where the displaced ligand is radiolabeled, and also allow for thedetermination of ligand-receptor binding affinity, based on competitivedisplacement of a binding agent whose binding affinity to the targetreceptor is known.

In both diagnostics and high-throughput screening, there is increasinginterest in developing electrochemical biosensors capable of detectingand quantifying ligand-receptor binding events. Such biosensors aredesigned to produce electrical signals in response to a selectedanalyte-specific event, such as a ligand-receptor binding event. Theinterest in biosensors is spurred by a number of potential advantagesover strictly biochemical assay formats, such as those discussed above.

First, biosensors may be produced, using conventional microchiptechnology, in highly reproducible and miniaturized form, with thecapability of placing a large number of biosensor elements on a singlesubstrate.

Secondly, because small electrochemical signals can be readily amplified(and subjected to various types of signal processing if desired),biosensors have the potential for measuring minute quantities ofanalyte, and proportionately small changes in analyte levels.

A consequence of the features above is that a large number of differentanalytes can be detected or quantitated by applying a small samplevolume, e.g., 10-50 μl, to a single multi-sensor chip.

Heretofore, electrochemical biosensors have been more successfullyapplied to detecting analytes that are themselves electrochemicalspecies, or can be participate in catalytic reactions that generateelectrochemical species, than to detecting ligand-receptor bindingevents. This is not surprising, given the more difficult challenge ofconverting a biochemical binding event to an electrochemical signal. Oneapproach to this problem is to provide two separate reaction elements inthe biosensor: a first element contains a receptor and boundenzyme-linked ligand, and the second element, components forenzymatically generating and then measuring an electrochemical species.In operation, analyte ligand displaces the ligand-enzyme conjugate fromthe first element, releasing the enzyme into the second element region,thus generating an electrochemical species which is measured in thesecond element.

Two-element biosensors of this type are relatively complicated toproduce, particularly by conventional silicon-wafer methods, since oneor more biological layers and permselective layers must be deposited aspart of the manufacturing process. Further, enzymes or receptors in thebiosensor can denature on storage, and the device may have variable"wetting" periods after a sample is applied.

Biosensors that attempt to couple electrochemical activity directly to aligand-receptor binding event, by means of gated membrane electrodes,have been proposed. For example, U.S. Pat. Nos. 5,204,239 and 5,368,712disclose gated membrane electrodes formed of a lipid bilayer membranecontaining an ion-channel receptor that is either opened or closed byligand binding to the receptor. Electrodes of this type are difficult tomake and store, and are limited at present to a rather small group ofreceptor proteins.

Alternatively, direct ligand/receptor binding may be measuredelectrically by embedding the receptor in a thin polymer film, andmeasuring changes in the film's electrical properties, e.g., impedance,due to ligand binding to the receptors. U.S. Pat. No. 5,192,507 isexemplary. Since ligand binding to the receptor will have a rather smalleffect on film properties, and since no amplification effect isachieved, the approach is expected to have limited sensitivity.

It would thus be desirable to provide a biosensor capable of detectingand quantifying ligand-binding events and characterized by: (i) directelectrochemical conversion of the binding event to electrical signal;(ii) a high electron flow "turnover" from each binding event; (iii)adaptable to substantially any ligand, and (iv) good storagecharacteristics and rapid wetting with sample application. In additionsthe device should be easily produced, and preferably amenable tomanufacture using standard microchip technologies.

SUMMARY OF THE INVENTION

One aspect of the invention is a biosensor apparatus for detecting abinding event between a ligand and ligand-binding receptor. An electrodein the apparatus includes an electrode substrate with a detectionsurface covered by a monolayer of hydrocarbon chains. The chains areanchored at their proximal ends to the detection surface, and aresufficiently close-packed and ordered to form an effective barrier toelectron flow across the monolayer mediated by a redox ion species in anaqueous solution in contact with the monolayer.

The ligand whose binding to a receptor is to be detected is attached tothe distal ends of a portion of the monolayer chains, such that bindingof a ligand-binding receptor to ligand perturbs the monolayersufficiently to measurably increase electron flow across the monolayermediated by such redox ion species.

The aqueous solution of redox species in contact with the monolayer isheld in a chamber that is also designed to receive sample receptor, tobring the receptor into contact with ligand on the monolayer.Ion-mediated electron flow across said monolayer, in response to bindingevents occurring between said receptor and ligand, is measured in anelectrical circuit in the apparatus.

In a preferred embodiment, the monolayer is composed of 8-22 carbon atomchains attached at their proximal ends to the detection surface, e.g., agold surface, by a thiolate linkage. The chains have a preferredmolecular density of about 3 to 5 chains/nm².

The dielectric constant of the monolayer in the presence of the solutionof redox species, but in the absence of the binding receptor, ispreferably less than about 2, with a change in the dielectric constantof 10% or more, by receptor binding to the ligand, being readilydetectable.

Exemplary ligand-receptor pairs include antigen-antibody,hormone-receptor, drug-receptor, cell-surface antigen-lectin,biotin-avidin, substrate/antibody and complementary nucleic acidstrands, where the ligand is typically the first-named of these pairs.Where the apparatus is used to detect a ligand or analog of the ligand,the apparatus may further include a receptor which competes with theanalyte ligand or analog for binding to the ligand on the monolayer. Oneexemplary ligand is an oligosaccharide ligand, and one exemplaryreceptor, the Verotoxin receptor, also known "Shiga-like toxin".

The electrode employed in the biosensor may be prepared, in accordancewith another aspect of the invention, by (i) subjecting the conductivemetal surface of the electrode substrate to mild oxidation conditions,(ii) adding to the substrate, a solution of hydrocarbon chains havinglengths between 8-22 carbon atoms and derivatized at one chain end witha thiol group, and (iii) applying a positive potential to the electrode.The potential placed on the electrode is preferably at least 250 mV vsNHE (normal hydrogen electrode), in a solution containing the alkylthiol to be deposited, and electrolytes including lithium ion andperchlorate anions. A selected portion of the hydrocarbon chains arederivatized at their ends opposite the thiol group, with the ligand ofinterest. The oxidative conditions applied to the electrode surface aresuch as to produce deposition of a monolayer of close-packed, orientedchains on the substrate, as evidenced by the ability of the electrode toform an effective barrier to electron ion flow across the monolayermediated by a redox ion species in an aqueous solution in contact withthe monolayer.

In another general embodiment of the biosensor apparatus, ligandmolecules are attached to the hydrocarbon chains forming the monolayerin the electrode through a heterodimer-subunit complex composed of firstand second peptides that together form α-helical coiled-coilheterodimer, where: (i) the first peptide is covalently bound to theelectrode surface through a spacer, such as an oligopeptide orhydrocarbon chain; (ii) the ligand is covalently attached to the secondpeptide; (iii) binding of the second peptide to the first peptide, toform such complex, is effective to measurably reduce the electron flowacross the monolayer mediated by such redox ion species, relative toelectron flow observed in the presence of the first peptide alone; and(iv) binding of a ligand-binding receptor to the ligand, with suchforming part of said complex, is effective to measurably increase theelectron flow across of the monolayer mediated by such redox species.

Also contemplated is an electrode for use in a biosensor apparatus ofthis type, composed of a substrate having a detection surface and ligandmolecules attached to surface through an α-helical coiled-coilheterodimer of the type detailed above.

The electrode just described can be produced, in accordance with anotheraspect of the invention, by contacting together: (a) a detection surfacehaving attached thereto, a first heterodimer-subunit peptide, and (b) asecond heterodimer subunit capable of binding to the first subunit toform an α-helical heterodimer, and having a covalently attached ligandcapable of binding specifically to such ligand-specific receptor.

These and other objects and features of the invention will become morefully apparent when the following detailed description of the inventionis read in conjunction with the accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a simplified, partly schematic view of the a biosensorapparatus constructed in accordance with the invention;

FIG. 2 is an enlarged view of a region the electrode in the biosensorshown in FIG. 1;

FIGS. 3A-3C illustrate three methods for forming a biosensor electrodehaving a lipid monolayer and attached ligand molecules, in accordancewith the invention;

FIG. 4. is a plot of monolayer thickness as a function of appliedvoltage in an electrode monolayer formed in accordance with the methodillustrated in FIG. 3B;

FIG. 5 illustrates the triggering of conductance by receptor-ligandinteraction on a biosensor electrode, in accordance with the invention;

FIGS. 6A and 6B illustrate the perturbation of lipid monolayer structurewith binding of PAK peptide to disaccharide ligands on a monolayer;

FIG. 7 shows plots of changes in oxidation (solid circles) and reduction(open squares) current of Fe(CN)₆ ³⁻ /⁴⁻ as a function of time afteraddition of PAK peptide to the monolayer illustrated in FIGS. 6A and 6B;

FIGS. 8A-8C illustrate the perturbation of lipid monolayer structurewith binding of Verotoxin to trisaccharide ligands on a monolayer;

FIG. 9 shows plots of changes in oxidation (solid circles) and reduction(open squares) current of Fe(CN)₆ ³⁻ /⁴⁻ as a function of time afteraddition of Verotoxin to the monolayer illustrated in FIGS. 8A and 8B;

FIG. 10 is a plot of electrode current of Fe(CN)₆ ³⁻ /⁴⁻ as a functionof temperature in a monolayer electrode constructed in accordance withthe invention;

FIGS. 11A and 11B demonstrate ion gating effects with a negativelycharged ligand in an electrode monolayer;

FIGS. 12A and 12B demonstrate ion gating effects with a positivelycharged ligand in an electrode monolayer;

FIGS. 13A and 13B illustrate the structure of an electrode monolayerhaving an embedded K coil peptide subunit (13A), and an embedded Kcoil/E coil heteroduplex;

FIG. 14 shows the change in oxidation (solid circles) and reduction(open squares) current as a function of time after addition of E coilpeptide subunit to an electrode of the type illustrated in FIG. 13Acontaining an embedded K coil peptide subunit;

FIG. 15 shows changes in oxidation of Fe(CN)₆ ³⁻ /⁴⁻ (open circles) andreduction (open squares) as a function of time after addition of PAKpeptide to an electrode containing di-saccharide ligands on a K coil/Ecoil lipid monolayer;

FIG. 16 shows changes in oxidation of Fe(CN)₆ ³⁻ /⁴⁻ (open circles) andreduction (open squares) as a function of time after addition ofVerotoxin peptide to an electrode containing trisaccharide ligands on aK coil/E coil lipid monolayer;

FIGS. 17A-17E show a synthetic pathway used for producing atrisaccharide-hydrocarbon conjugate employed in the monolayer shown inFIGS. 8A-8C; and

FIG. 18 shows a synthetic pathway used in producing adisaccharide-hydrocarbon conjugate employed in the monolayer shown inFIGS. 6A and 6B.

DETAILED DESCRIPTION OF THE INVENTION

A. Biosensor Apparatus

FIG. 1 is a simplified schematic view of a biosensor apparatus 20 fordetecting a binding event between a ligand and a ligand-binding receptoror agent, in accordance with the invention. The apparatus includes aworking electrode 22 having a conductive detection surface 24, and ahydrocarbon-chain monolayer 26 formed on the detection surface. In theembodiment shown, the detection surface is the upper surface of aconductive film 28 deposited on an electrode substrate 30, which may benon-conductive material. Details of the monolayer formed on thedetection surface, and the method of forming the monolayer on thesurface, are discussed below.

A cover 32 in the apparatus has an upper wall 34, and side walls, suchas wall 36, which are joined to edge regions of the electrode substrateto form a closed chamber 38 therewith. The chamber serves to hold anaqueous electrolyte solution required for biosensor operation, as willbe described. Liquid may be introduced into or withdrawn from thechamber through a valved port 39 as shown. Although not shown, thechamber may include a second port or vent to facilitate liquid flowthrough the port.

A reference electrode 40 and a counter electrode 42 in the apparatus arecarried on the chamber-facing surface of wall 34, as shown, and are thusboth in conductive contact with electrode 22 when the chamber is filledwith electrolyte solution. The reference electrode, which is held atground, serves as the voltage potential reference of the workingelectrode, when a selected potential is placed on the working electrodeby a voltage source 44. This potential is measured by a voltagemeasuring device 46 which may additionally include conventionalcircuitry for maintaining the potential at a selected voltage, typicallybetween about -500 to +800 mV.

Voltage source 44 is connected to counter electrode 42 through a currentmeasuring device 48 as shown, for measuring current flow between the twoelectrodes during biosensor operation. The reference and counterelectrodes are Pt, Ag, Ag/AgCl, or other suitable electrodes. Thereference and counter electrodes, and the circuitry connecting them tothe working electrode, are also referred to herein, collectively, asmeans for measuring ion-mediated electron flow across theworking-electrode monolayer, in response to ligand-receptor bindingevents occurring at the monolayer surface.

FIG. 2 is an enlarged view of a portion of the working electrode,including the electrode monolayer, showing individual hydrocarbonchains, such as chains 50, forming the monolayer, and ligand molecules,such as molecules 52, covalently attached to distal ends of thehydrocarbon chains. The ligand employed in the biosensor is a selectedbinding partner in a ligand/receptor binding pair, where the analyte tobe detected is related to one of the two binding partners.Ligand-receptor binding pairs used commonly in diagnostics includeantigen-antibody, hormone-receptor, drug-receptor, cell surfaceantigen-lectin, biotin-avidin, and complementary nucleic acid strands,where the ligand is typically the smaller of the two binding pairmembers. The analyte to be detected may be either member of the bindingpair, or alternatively, a ligand analog that competes with the ligandfor binding to the complement receptor.

The ligand molecules are attached to distal ends of the chains throughconventional derivatization reactions, e.g., ester, ether, amide, orsulfhydryl linkages, according to standard methods. The number of chainsin the monolayer carrying distal-end ligands is preferably about 1 to 10mole percent of the total chains, but may range from 0.01 to 100%.

The chains forming the monolayer are typically 8-22 carbon, saturatedhydrocarbon chains, although longer chains, chains with someunsaturation, chains with noncarbon chain atoms, such as lipid ethers,and/or chains with minor branching, such as by non-chain methyl groups,may be employed, within the constraint that the chains, at a sufficientpacking density, form a sufficiently close packed and ordered monolayerto be effective as a barrier to electron flow, under biosensor operatingconditions, as discussed below. This density is calculated to be between3-5 chains/nm².

As an example of the variation in chain composition allowed, theembodiment of the invention shown in FIG. 13B has a hydrocarbon-chainmonolayer that includes coil-coil peptide heterodimers embedded in theplanar chain matrix, while still retaining a low dielectric barrier toion flow through the monolayer.

In the embodiment shown, the chains are coupled to the electrodedetecting surface through sulfhydryl linkages, although other suitablecoupling groups may be employed. Methods for producing monolayers havingsuitable hydrocarbon chain densities will now be discussed.

B. Electrode Monolaver Production

FIGS. 3A-3C illustrate three methods for forming hydrocarbon chainmonolayers suitable for use in the biosensor electrodes.

One approach, illustrated in FIG. 3A, involves passive diffusion ofchains, such as hydrocarbon chains 54 and ligand-derivatized chains,such as chains 56, onto the surface an electrode 58, under conditionseffective to couple the diffused chains to the electrode detectionsurface. The diffusion method illustrated in 3A is a two-step process.In the first step, hydrocarbon chains alone (in the absence ofligand-derivatized chains) are allowed to react with the detectedsurface over an extended period, e.g., 24-48 hours, until a selectedpacking density less than full packing density is achieved.

The diffusion reaction is carried out under conditions suitable forcoupling the derivatized chains to the detection surface. Where thechains have thiol coupling groups, and the electrode surface is gold,the surface is subjected to mild electro-chemically oxidizingconditions, with a perchlorate salt present in solution, then reactedwith the chains under mildly oxidizing conditions.

The extent of packing can be monitored, for example, by ellipsometrymeasurements to determine the thickness of the layer on the detectionsurface. At maximum density, i.e., saturation, a given chain length willproduce a given monolayer thickness. As a guide, C₂₂ chains produce amaximum monolayer thickness of about 30 Å, and shorter length chains,proportionately thinner monolayers. Thus, in the case of a monolayerformed of C₂₂ chains, the passive buildup of the monolayer may bestopped when a 25 Å monolayer thickness is observed.

The second diffusion step involves the passive diffusion ofligand-derivatized thiol-chains 56 onto the partially formed monolayer,indicated at 60, again under suitable thiolate coupling conditions,until a high-density monolayer 62 is achieved, as evidenced, forexample, by the measured thickness of the monolayer and/or a plateauingof the thickness/time curve.

Although this approach has been applied successfully to monolayerproduction in the invention, it suffers from two limitations. First,rather long diffusion times--on the order of one to several days--arerequired to reach maximum packing density. Secondly, the percent chainscontaining attached ligands is difficult to control reproducibly, sothat the final monolayers will have variable mole percentages ofligands, and thus, different performance characteristics.

These limitations are substantially overcome in the method illustratedin FIG. 3B, in accordance with another novel aspect of the invention. Inthis approach, a mixture of free and ligand-carrying hydrocarbon chains,such as chains 66, 68, respectively, at a desired mole ratio, areactively driven to the surface by applying a positive voltage potentialto the substrate, here indicated at 64. In practice, the hydrocarbonchain mixture (about 1 mM hydrocarbon chains) in an ethanolic solutionof 100 mM Li perchlorate, neutral pH, is added placed over theelectrode, and a selected potential is applied to the electrode. Thebuildup of the monolayer can be monitored by increase in layerthickness, as above. Preferably, however, monolayer formation ismonitored by measuring electron flow across the monolayer, e.g.,employing the circuit configuration shown in FIG. 1. In this case,formation of the monolayer, indicated at 70, will be characterized by asteady drop in electrode current, until a stable low current flow isreached, at which point maximum chain packing has been achieved.

The time required to achieve saturation packing density will vary withapplied voltage, and can be a short as 10 seconds--that is, about 4orders of magnitude faster than monolayer formation by diffusion. FIG. 4is a plot of monolayer thickness formed using a thiol-group C₂₂hydrocarbon chain under coupling conditions like those above, after 10minutes at the electrode voltage indicated. As seen, complete or nearlycomplete monolayer formation (30 Å thickness) occurs within 10 minutesat about 1 V (vs. NHE) potential and above. At lower positive voltages,additional reaction time is required. Preferably the voltage applied tothe electrode is at least voltage between about +250 mV relative to anormal hydrogen electrode (+250 vs. NHE) and 1.2 V (vs. NHE).

Not only are rapid monolayer formation times achieved, but thepercentages of ligand- and non-ligand chains present in the reactionmixture are precisely represented in the monolayers, giving highlyreproducible electrode characteristics.

FIG. 3C shows a modification of the FIG. 3B method, where thehydrocarbon-chain mixture reacted with the electrode (indicated at 71)includes non-ligand chains, such as chains 72, and peptide subunitconjugates, such as indicated at 74, containing a peptide subunit 76that is capable of forming a stabilized, alpha-helical peptideheterodimer with an oppositely charged, complementary subunit. Suchheterodimer subunits are described in PCT patent application WO CA95/00293, for "Heterodimer Polypeptide Immunogen Carrier Composition andMethod", publication date Nov. 23, 1995, which is incorporated herein byreference. Exemplary subunits are referred to herein as K coils,referring to a positively charged subunits whose charge is provideddominantly by lysine residues, and E coils, referring to negativelysubunits whose charge is provided dominantly by glutamic acid residues.

In the embodiment shown, subunit 76 is attached to the distal end of ahydrocarbon chain 78 (end opposite the chain's thiol group) by suitablelipid-to-peptide conjugation, e.g., by ester linkage to a hydrocarbonfatty acid. Alternatively, and as described below, the peptide subunitmay be linked to the electrode surface through a peptide spacer, e.g.,tripeptide spacer that extends from one end of the subunit and includescysteine as a terminal residue, for sulfhydryl attachment to theelectrode surface. In both cases, the peptide subunit conjugate is mixedwith the hydrocarbon chains, at a selected mole ratio, then driven intoa monolayer formation by applying a positive voltage to the electrode,as above, until a densely packed monolayer 80 is formed.

A suitable ligand is then attached to the monolayer by contacting themonolayer with a ligand-coil conjugate 82 composed of the oppositelycharged complement of the monolayer coil, indicated at 84, coupled to aselected ligand 86. The two oppositely charged subunits spontaneouslyself-assemble into heterodimers, effectively coupling the ligand to themonolayers with the high affinity constant of the two heterodimers.

The method provides, in addition to the advantages mentioned above withrespect to FIG. 3B, a "universal" biosensor substrate which can bemodified to include one of a large number of different ligands in thesubstrate monolayer, simply by contacting the universal substrate with aconjugate of the oppositely charged peptide subunit and the selectedligand. In the example shown in FIG. 3C, a universal substrate monolayer80 is converted to a ligand-specific monolayer 88 by addition of theligand-specific conjugate 82.

C. Biosensor Characteristics: Directly Attached Ligand

This section examines the dielectric properties of the biosensors of theinvention, as evidenced by the conductance properties of the biosensormonolayer membranes in the presence and absence of ligand-receptorbinding. The present section considers membranes having directlyattached ligands of the type described with respect to FIGS. 3A and 3B.The next section examines similar electrical properties in biosensormembranes in which the ligand is attached through heterodimer peptidesubunits, as described with respect to FIG. 3C.

The basic operational features of the biosensor are illustrated in FIG.5. The figure shows a biosensor electrode 90 in a biosensor apparatus ofthe type described in FIG. 1, where an electrode monolayer 92 is formed,as above, of a densely ordered array of hydrocarbon chains containingligand molecules, such as molecule 94, attached to the distal ends ofsome of the chains.

The electrode is in contact with a solution of ionic species, indicatedat 98, capable of undergoing a redox reaction, i.e., losing or gainingan electron, at a suitably charged electrode. Exemplary redox speciesare Fe(CN)₆ ³⁻ /⁴⁻ as a negatively charged species, and Ru(NH₃)₆^(2+/3-) as a positively charged species. Other probes which can be usedinclude Mo(CN)₆ ³⁻ (E₀ =+800 mV), W(CN) ₆ ³⁻ (E₀ =+580 mV), Fe(CN)₄ ⁻(E₀ =+580 mV), Ce^(4+/3+), (E₀ =+1.4 V), and Fe^(+3/2+) (E₀ =+666 mV).Typical redox ion concentrations are between 0.01 and 10 mM. The redoxsolution is contained in chamber, like chamber 38 in FIG. 1, and is incontact with reference and counter electrodes.

The voltage potential placed on the electrode, i.e., between theelectrode and reference electrode, is typically at least 90 mV above theelectrochemical potential (e₀) value of the redox species, foroxidation, and at least 90 mV below the electrochemical potential, forreduction of the species. Consider, for example, Fe(CN)₆ ^(3-/4-), withan E₀ of 450 mV (vs. NHE). Above about 550 mV electrode potential, anyFe2+ species is oxidized to Fe3+, and at an electrode potential belowabout 350 mV, and Fe+3 is reduced to Fe+2. Similarly, Ru(NH₃)₆ ^(2+/3+)has an E₀ of +50 mV (vs. NHE), so oxidation is achieved at an electrodepotential above about +150 mV, and reduction, below about -50 mV.

In the absence of receptor binding to the ligand, the monolayer retainsits dense ordered packing, forming an effective barrier to electron flowacross the monolayer mediated by the redox ion species, when a suitableoxidizing or reducing potential is placed across the monolayer. This isreflected by a low or zero measured current across the membrane. Thedielectric constant of the monolayer in this condition is typicallyabout 1-2.

With binding of a receptor 96 to a ligand on a monolayer, as shown atthe right in the figure, the ordered structure of the monolayer isperturbed sufficiently to allow the movement of redox species throughthe monolayer, producing electron flow through the electrode.Measurements performed in support of the invention indicate that onetriggering event leads to 10² to 10⁶ ionic and electron transfer eventsper second, and thus is highly multiplicative. The biosensor recordsthis binding event as an increase in current across the electrode, i.e.,between the working and counter electrodes.

By analogy to a transistor, the redox solution serves as the "source",the monolayer as the "gate", and the underlying electrode as the"drain". Current flow in a transistor is initiated by applying athreshold voltage to the gate. In the biosensor of the invention,current flow is initiated by a stimulus--in this case, a ligand-receptorbinding event--to the monolayer "gate".

A biosensor electrode 100 constructed in accordance with the invention,and having a disaccharide ligand indicated at 102 is shown before andafter receptor binding in FIGS. 6A and 6B, respectively. Synthesis ofthe disaccharide-hydrocarbon chain used in the membrane is described inExamples 1D and 1E. The electrode was prepared as described withreference to FIG. 3B, employing a ratio of non-ligand to ligand-chainsof about 4 to 1. The disaccharide is specifically reactive with aPseudomonas PAK peptide, indicated at 104, forming a ligand-receptorpair with the peptide.

The increase in biosensor electrode current, when PAK peptide receptoris added to the biosensor chamber, is seen in FIG. 7 for both oxidation(solid circles) and reduction (open squares) current from Fe(CN)₆ ³⁻/⁴⁻. The increase over time presumably reflects the kinetics of binding,demonstrating that the biosensor is useful as well in measuring the rateligand-receptor binding events. FIG. 6B illustrates the perturbation ofthe hydrocarbon chain structure with receptor binding.

As another example, the biosensor electrode illustrated in FIGS. 8A-8C(electrode 22 from FIG. 2) has a trisaccharide ligand 52 which is shownbefore and after receptor binding in FIG. 8A and FIGS. 8B and 8C,respectively. Synthesis of the trisaccharide-hydrocarbon chain used inthe membrane is described in Examples 1B and 1C. The electrode wasprepared as described with reference to FIG. 3B, employing a ratio ofnon-ligand to ligand-chains of about 4 to 1. The disaccharide isspecifically reactive with a Verotoxin, indicated at 106, forming aligand-receptor pair. Verotoxin was prepared as described in Example 2.

FIGS. 8B and 8C illustrate two possible binding configurations. Theconfiguration in FIG. 8B has little effect on the monolayer structure,and hence on biosensor current, because binding is "remote" from themembrane surface; the configuration illustrated in FIG. 8C, by contrast,produces significant perturbation of the monolayer structure, and thuswould be expected to significantly enhance biosensor current.

The oxidation and reduction current plots shown in FIG. 9 demonstratethat Verotoxin binding to the membrane does in fact produce a majorchange in monolayer structure. As seen, both oxidation and reductioncurrent increase from near-zero levels, in the absence of Verotoxin, toa level in the μAmp range an hour after Verotoxin is introduced into thebiosensor.

In the examples above, the stimulation of biosensor current by receptorbinding may be the result of (i) steric perturbation of the monolayerchains, as indicated in FIGS. 6B and 8C, (ii) charge effects on themonolayer surface due to charged groups on the receptor, or (iii) acombination of the both effects. Studies conducted in support of theinvention indicate that both effects can be operative.

The effect of hydrocarbon-chain disruption in the biosensor monolayer,was examined by plotting biosensor current as a function of electrodetemperature. If lipid-chain disruption leads to greater electron flow inthe biosensor, raising the temperature of the monolayer, and thus themotion of the lipid chains, should increase measured electron flowmediated by redox carriers. This was in fact observed, as seen in FIG.10. The current/temperature plot has a peak corresponding to the phasetransition temperature of the monolayer chains (about 55° C.),consistent with the idea that maximum lipid disruption occurs at thepoint of maximum extent of phase boundaries in the hydrocarbon chains.

The effect on conductance of charge on the monolayer surface can be seenfrom FIGS. 11 and 12. In the study represented in FIG. 11A, a negativelycharged ligand was attached to the distal ends of a portion of thechains forming the monolayer. In the figure, the electrode is indicatedat 108, the monolayer, at 110, chains forming the monolayer, at 112, andchain-attached ligands, at 114. Electrode current was measured for thenegatively charged redox species Fe(CN)₆ ^(3-/4-), and independently,with the positively charged species Ru(NH₃)₆ ^(2+/3+), at oxidationpotentials indicated above.

As seen in FIG. 11B, the oxidation current for the positively chargedspecies shows the ion-dependent behavior expected for ion migrationthrough the monolayer, indicating that the monolayer is conductive topositively charged redox species. Conversely, no significant electronflow was observed with the negatively charged redox species.

Similar results were obtained with a monolayer designed to contain apositively charged surface ligand, as illustrated in FIG. 12A. In thisfigure, the electrode is indicated at 116, the monolayer, at 118, chainsforming the monolayer, at 120, and chain-attached ligands, at 122. Asseen in FIG. 12B, ion-dependent current was observed for oxidation ofthe negatively charged iron redox species, but not the positivelycharged ruthenium species.

D. Biosensor Characteristics: Heterodimer Attached Ligand

In another embodiment, the ligand in the biosensor is anchored tobiosensor surface, i.e., embedded within the hydrocarbon-chainmonolayer, by a coiled-coil heterodimer complex formed of two subunitpeptides. The heterodimer-subunit peptides employed in the biosensorinvention are two non-identical, preferably oppositely chargedpolypeptide chains, typically each about 21 to about 70 residues inlength, having an amino acid sequence compatible with their formationinto two-stranded α-helical heterodimeric coiled-coils. They aredesignated herein as HSP1 (heterodimer-subunit peptide 1), and HSP2(heterodimer-subunit peptide 2). In the discussion below, HSP1 willrefer to the peptide attached to the biosensor surface in the biosensor,and HSP2, to the peptide having an attached ligand. It will beunderstood that these designations refer to the functional role playedby the subunit peptide, not the actual peptide sequence.

In aqueous medium, the isolated heterodimer-subunit peptides aretypically random coils. When HSP1 and HSP2 are mixed together underconditions favoring the formation of α-helical coiled-coil heterodimers,they interact to form a two-subunit α-helical coiled-coil heterodimericcomplex.

Peptides in an α-helical coiled-coil conformation interact with oneanother in a characteristic manner that is determined by the primarysequence of each peptide: The tertiary structure of an α-helix is suchthat 7 amino acid residues in the primary sequence correspond toapproximately 2 turns of the α-helix. Accordingly, a primary amino acidsequence giving rise to an α-helical conformation may be broken downinto units of 7 residues each, termed heptads. The heterodimer-subunitpeptides are composed of a series of heptads in tandem. When thesequence of a heptad is repeated in a particular heterodimer-subunitpeptide, the heptad may be referred to as a "heptad repeat", or simply"repeat".

Specific types of amino acid residues at defined positions in eachheptad act to stabilize the two-stranded α-helical coiled-coilheterodimeric structure or complex. The heterodimer peptides may alsocontain residues that can be reacted (either intra- or interhelically)to stabilize the α-helical or coiled-coil nature of the polypeptides.One example of a stabilizing modification is the incorporation of lactambridges in the first and last (terminal) repeats of heterodimer-subunitpeptides, as detailed in PCT application Wo CA95/00293 for "HeterodimerPolypeptide Immunogen Carrier Composition and Method", publication dateNov. 23, 1995, which is incorporated herein by reference.

The dimerization of HSP1 and HSP2 is due to the presence of a repeatedheptad motif of conserved amino acid residues in each peptide's primaryamino acid sequence. Repeating heptad motifs having appropriate aminoacid sequences direct the HSP1 and HSP2 polypeptides to assemble into aheterodimeric α-helical coiled-coil structure under permissibleconditions. The individual α-helical peptides contact one another alongtheir respective hydrophobic faces.

HSP1 and HSP2 may assemble into a heterodimer coiled-coil helix(coiled-coil heterodimer) in either parallel or antiparallelconfigurations. In a parallel configuration, the two heterodimer-subunitpeptide helixes are aligned such that they have the same orientation(amino-terminal to carboxyl-terminal). In an antiparallel configuration,the helixes are arranged such that the amino-terminal end of one helixis aligned with the carboxyl-terminal end of the other helix, and viceversa.

Heterodimer-subunit peptides designed in accord with the guidancepresented in the above PCT application typically show a preference forassembling in a parallel orientation vs. an antiparallel orientation.For example, the exemplary peptides identified by SEQ ID NO:1 and SEQ IDNO:2 in the above CA95/00293 PCT patent application, formparallel-configuration heterodimers as do other peptide sequencesdiscussed in the PCT application. When attaching a ligand to HSP2, it isgenerally desirable to attach the ligand at or near the end of thepeptide that will form the distal end of the heterodimer. In particular,where the heterodimer forms a parallel configuration, the HSP1 peptideis preferably anchored to the biosensor surface at its C terminus, andthe ligand attached to the HSP2 peptide at its N terminus.

As just noted, one of the two subunit peptides (HSP1) in the heterodimeris attached to the biosensor surface, and the second peptide (HSP2)contains a ligand intended to participate in an analyte-dependentligand/anti-ligand binding reaction. In both cases, the peptide issynthesized, or derivatized after synthesis, to provide the requisiteattachment function and ligand, respectively.

Considering the modification of HSP1, the peptide may be synthesized, ateither its N or C terminus, to carry additional terminal peptides thatcan function as a spacer between the biosensor surface and thehelical-forming part of the peptide. Alternatively, the HSP1 peptide canbe attached to the biosensor surface thorough a high-affinity bindingreaction, such as between a biotin moiety carried on the peptide and anavidin molecule covalently attached to the surface.

Where the heterodimer is embedded in a hydrocarbonchain monolayer, asdescribed below, the spacer anchoring the HSP1 peptide to the biosensorsurface may be a hydrocarbon chain. The chain is preferably a fractionallength of the chains making up the bilayer, such that the distal ends ofthe heterodimer peptides in the assembled monolayer are at or near theexposed surface of the monolayer. Thus, for example, if the monolayer ismade up of 18-carbon chains, the spacer is preferably 2-10 carbons inlength, depending on the length of the assembled heterodimer.

The hydrocarbon-chain spacer, in the form of a omega-thio fatty acid,may be coupled to a terminal hydroxyl or amine coupling duringsolid-phase synthesis, as outlined above. The derivatized peptide, inturn, can be attached to a metal surface by standard thiolate coupling(Dakkouri, et al., Langmuir (1996) 12:2849-2852).

Considering the ligand-attachment to HSP2, the ligand selected will bedetermined by the analyte to be tested. Ligand-receptor binding pairs,i.e., ligand/ligand-binding agent pairs used commonly in diagnosticsinclude antigen-antibody, hormone-receptor, drug-receptor, cell surfaceantigen-lectin, biotin-avidin, substrate/enzyme, and complementarynucleic acid strands. The ligand is typically the smaller of the twobinding pair members, particularly where the ligand is attached to ahydrocarbon-chain monolayer, as described below. However, attachment ofeither binding pair is contemplated herein.

Where the ligand is a polypeptide, e.g., peptide antigen, the antigencan be synthesized by either solid-state or recombinant methods, toinclude the peptide antigen at the end of the HSP2 peptide that willorient distally in the assembled heterodimer. Where the ligand is anon-peptide moiety, e.g., a non-peptide hormone, drug, or nucleic acid,the HSP2 peptide can be synthesized to include one or more residues thatcan be specifically derivatized with the ligand. The ligand ispreferably covalently attached to the N-terminal amino acid residue, orto one or the residues facing the exposed face of the heterodimer.Preferred coupling groups are the thiol groups of cysteine residues,which are easily modified by standard methods. Other useful couplinggroups include the thioester of methionine, the imidazolyl group ofhistidine, the guanidinyl group of arginine, the phenolic group oftyrosine and the indolyl group of tryptophan. These coupling groups canbe derivatized using reaction conditions known to those skilled in theart.

To attach the ligand-derivatized HSP2 peptide to the surface-immobilizedHSP1 peptide, the two peptides are contacted under conditions that favorheterodimer formation. A medium favoring coiled-coil heterodimerformation is a physiologically-compatible aqueous solution typicallyhaving a pH of between about 6 and about 8 and a salt concentration ofbetween about 50 mM and about 500 mM. Preferably, the salt concentrationis between about 100 mM and about 200 mM. An exemplary benign medium hasthe following composition: 50 mM potassium phosphate, 100 mM KCl, pH 7.Equally effective media may be made by substituting, for example, sodiumphosphate for potassium phosphate and/or NaCl for KCl. Heterodimers mayform under conditions outside the above pH and salt range, medium, butsome of the molecular interactions and relative stability ofheterodimers vs. homodimers may differ from characteristics detailedabove. For example, ionic interactions between the ionic groups thattend to stabilize heterodimers may break down at low or high pH valuesdue to the protonation of, for example, Glu side chains at acidic pH, orthe deprotonation of, for example, Lys side chains at basic pH. Sucheffects of low and high pH values on coiled-coil heterodimer formationmay be overcome, however, by increasing salt concentration.

Increasing the salt concentration can neutralize the stabilizing ionicattractions or suppress the destabilizing ionic repulsions. Certainsalts have greater efficacy at neutralizing the ionic interactions. Forexample, in the case of the K-coil peptide in FIG. 13A, a 1M or greaterconcentration of ClO₄ ⁻ anions is required to induce maximal α-helicalstructure, whereas a 3M or greater concentration of Cl⁻ ions is requiredfor the same effect. The effects of high salt on coiled-coil formationat low and high pH also show that interhelical ionic attractions are notessential for helix formation, but rather, control whether a coiled-coiltends to form as a heterodimer vs. a homodimer.

FIG. 13A shows a biosensor electrode 124 in which the hydrocarbon chainmonolayer, indicated at 126 includes a K coil peptide subunits, such assubunit 128, as described above. In the embodiment shown, each peptidesubunit is coupled to the electrode surface via a tripeptide spacer,such as spacer 127 in subunit 128, which is itself attached to theelectrode surface through a sulfhydryl linkage, as shown. The peptide,including the peptide spacer, is formed conventionally, e.g., by solidphase synthesis. The amount of peptide subunit in the monolayer is about20 mole percent. The monolayer was formed according to the methoddescribed above with respect to FIG. 3C. As indicated above, the peptidesubunit may alternatively be coupled to the distal ends of a portion ofthe hydrocarbon chains in the monolayer, placing the subunit more on themonolayer surface. A hydrocarbon chain-peptide conjugate suitable forthis application may be made, for example, by attaching an activated-endhydrocarbon chain to the terminal amino acid of the peptide, as theterminal step in solid phase synthesis.

Presumably because of the positive charge imparted to the monolayer bythe K coil subunits, the monolayer shows relatively high conductance tonegatively charged redox species, such as Fe(CN)₆ ³⁻, as evidenced by arelatively high oxidation or reduction current with the redox species.

FIG. 13B shows the same monolayer, but after addition of complementary,negatively charged E coil subunits, such as indicated at 130. As shown,oppositely charged subunits pair to form charge-neutral heterodimers inthe monolayer. This pairing is effective to reduce monolayer conductancesubstantially, as evidenced by the time-dependent fall in measuredoxidation or reduction current in the presence of Fe(CN)₆ ³⁻ ions (FIG.14).

As shown in FIG. 3C, the second peptide subunit, e.g., the E coilsubunit, added to the monolayer may be derivatized with a ligand,producing a monolayer having charge-neutral heterodimers embeddedtherein (or attached to the monolayer surface), and a ligand exposed onthe monolayer surface. The resulting electrode is effective to measureligand-specific receptor binding events in a biosensor operated inaccordance with the invention.

The operating characteristics of such a biosensor are illustrated inFIG. 15. The electrode in this biosensor includes (i) a monolayer withembedded K coils covalently attached to the electrode surface, (ii)complementary E coils forming heterodimers with the K coils in themonolayer, and (iii) surface disaccharide ligands of the type shown inFIG. 6 attached covalently to the E coils and disposed therefore at themonolayer surface. As seen in FIG. 15, addition of the PAK proteinreceptor (see FIG. 6B) produces a increase in both oxidation andreduction currents, with the current increase over time presumablyreflecting additional binding events after receptor addition to thebiosensor electrode.

A similar biosensor having a trisaccharide, rather than disaccharide,ligand attached to the E coil subunit in the electrode monolayer wastested with the Verotoxin receptor described above with respect to FIGS.8A-8C, with the results seen in FIG. 16. The solid lines in the figureshow the increase in oxidation and reduction current observed, as afunction of time, after addition of Verotoxin.

The following examples are intended to illustrate, but in no way limitthe invention.

EXAMPLE 1 Synthesis of Receptors in a Form Suitable for Immobilizationon a Gold Electrode

Selective tosylation of 1,16-dihydroxyhexane provided the monotosylatedalcohol 1 in 42% yield. Trisaccharide 2, obtained as described in theliterature (Janson, et al., J. Org. Chem. 53:5629 (1988)), was convertedinto an anomeric mixture of trichloroacetamidates 3. Glycosylation ofalcohol 1 with glycosyl donor 3 in CH₂ Cl₂ in the presence of acatalytic amount of trimethylsilyl trifluoromethanesulfonate gavetrisaccharide glycoside ω-tosylate 4, which was used in the next stepwithout purification. The tosyloxy group of compound 4 was displaced bythiocyanate to provide the trisaccharide glycoside 5, terminated at thereducing end by spacer-arm containing the masked thiol function.Reduction of thiocyanate by the action of sodium borohydride (Olsen, R.K., and Snyder, H. R., J. Org. Chem. 30:184 (1965).) followed bysaponification of acetate groups gave trisaccharide receptor 6.

The disaccharide imidate 7 was reacted with alcohol 1 in a similarfashion to that described for the trisaccharide 4. Synthesis of thedisaccharide glycosyl donor 7 is not described here but followsestablished methods that are considered a general art. Nucleophilicsubstitution of the tosyloxy group by thiocyanate was carried out asdescribed for preparation of 5 to give compound 8. Reduction ofthiocyanate accompanies by deacetylation afforded synthetic disaccharidereceptor 9.

A. 16-(p-Toluensulfonyloxy)hexadecanol (Structure 1)

To a solution of 1.1 g of 1,16-dihydroxyhexadecane in 10 ml of drypyridine 0.8 g of tosyl chloride was added. After 2 h mixture wasconcentrated diluted with 20 ml of acetone, 5 g of SiO₂ was added andacetone was removed in vacuum. The solid was applied on SiO₂ and elutedwith pentane-ethyl acetate (2:1) to yield 748 mg (42%) of C-101.

B.2,3,4,6-tetra-O-acetyl-D-galactopyranasyl(α1→4)-6-O-acetyl-2,3-di-O-benzoyl-D-galactopyranosyl(β1→4)-2,3,6-tri-O-benzoyl-D-glucopyranosyl(β1→O)-(16-thiocyano)hexadecanol(Structure 5)

A mixture of 277 mg of imidate, 100 mg of C-101 and 0.5 g of mol. sieves(4A) was stirred for 1 h. Then 8 μl of TMSOTf was added. After 2 h 1 mlof EA was added, solid was removed by filtration. Filtrate wasconcentrated and dried in vacuum. A solution of the residue and 200 mgof KSCN in 6 ml of DMF was stirred at 80° C. for 2 hours. Mixture wasconcentrated, dissolved in 30 ml of CH₂ Cl₂, washed with water andconcentrated again. Chromatography of the residue on SiO₂ withpentane-ethyl acetate (3:2) gave 225 ml (73%) of C-105.

C. D-Galactopyranosyl(α1→4)-D-galactopyranosyl(β1→4)-D-glucopyranosyl(β1→O)-(16-thio)hexadecanol(Structure 6)

To a solution of 60 mg of C-105 in 4 ml of dry MeOH ˜40 mg of NaBH₄ wasadded Ar. After stirring for 2 h at 45° C. mixture was concentrated anddissolved under gentle reflux in a solution of 50 mg of NaOH in 10 ml ofwater. After stirring overnight at 45° C., the mixture was neutralizedwith cationite (H⁺ -form) and applied on extraction medium SEPPAK(C-18). The cartridge was washed with 20, 40, 50, 80, and 100% solutionof MeOH in water. Fractions of 100% MeOH was concentrated to give 24.4mg (81%) of C-106.

D. (16-thiocyano)hexadecanyl4-O-(2-acetamido-3,4,6-tri-O-acetyl-2-D-galactopyranosyl)-2,3,6-O-acetyl-.beta.-D-galactopyranosideC-108 (Structure 8)

Mixture of 100 mg of imidate 7, 68 mg of C-101 and 100 mg of mol. sieves(4A) in 5 ml of CH₂ Cl₂ was stirred for 1 h. Then 5 μl of TMSOTf wasadded. After 2 h 1 ml of EA was added, solid was removed by filtration.Filtrate was concentrated and dried in vacuum. A solution of the residueand 70 mg of KSCN in 3 ml of DMF was stirred at 80° C. for 2 h. Mixturewas concentrated, dissolved in 30 ml of CH₂ Cl₂, washed with water andconcentrated again. Chromatography of the residue on SiO₂ withpentane-ethyl acetate (2:1) gave 82 mg (70%) of C-108.

E. (16-thiohydroxy)hexadecanyl4-O-(2-acetamido-2-deoxy-β-D-galactopyranosyl)-β-D-galactopyranoside(Structure 9)

A solution of 54.2 mg of 8 (C-108) and 40 mg of NaBH₄ in 3 ml of dryMeOH was refluxed for 2 h then neutralized with ion-exchange resin DOWEX(H⁺), concentrated and chromatographed on C-18 in H₂ O (50:50):100) toyield 19.7 mg (52%) of C-111.

EXAMPLE 2 Isolation of Verotoxin Receptor

Shiga-like (Vero) toxin I (SLT-I) was purified from Escherichia coliJM101 (pJB128) in a simple, one step procedure using solid-phaseextraction medium CHROMOSORB-P containing covalently coupled syntheticanalogs of the toxin's αGal(1,4)βGal (digalactoside) host cell receptors(SYNSORB-P1). Bacteria were grown in baffled Fernbach flasks at 37° C.in carbenicillin- (50 μg/ml) supplemented tryptic soy broth (TSB)containing SYNSORB-P1 (15 g/L) to bind toxin released from growingcells. Late log phase cultures were treated for 30 minutes at 37° C.with Polymyxin B sulfate (0.1 mg/mL) to release intracellular SLT-I andalso allow this to bind to the SYNSORB-P1. Next, the SYNSORB-P1 wascollected and washed thoroughly with 250 mM NaCl (pH 3.8) to removecells and cellular debris. THe SLT-I was eluted from the washedSYNSORB-P1 using 50 mM Tris base (pH 10) containing 250 mM NaCl (TN) andconcentrated using an Amicon ultrafiltration unit. The concentratedSLT-I was stable for weeks at 4° C. and could be frozen for extendedperiods of time without appreciable loss of activity. On average, 61%(n=10, SD mean=8, range 48% to 76%) of the SLT activity in the originalPolymyxin-treated TSB cultures was recovered in the TN fraction elutedfrom the SYNSORB-P1. SDS-polyacrylamide gel electrophoretic analysis ofthe SLT-I preparation revealed two prominent Coomassie blue-stainedbands. The molecular weight of these two bands was calculated to be35,000 and 7,500, respectively. The 7.5 KDz band reacted in westernimmunoblots with SLT-I but not SLT-II B subunit-specific monoclonalantibody. Amino terminal microsequence analysis of both bands confirmedtheir identity as the A and B subunits of SLT-I. Average yield of SLT-Iwas 0.32 mg/L (n=8, SD mean=0.3, range 0.1 to 0.8) of TSB culture andits specific activity in the Vero cytotoxicity assay was 4.4 pg/mL/CD₅₀.The results demonstrate the utility of SYNSORB in the facile and rapidpurification of carbohydrate binding toxins or lectins.

Although the invention has been described with respect to variousspecific embodiments and methods, it will be appreciated that variousmodifications and changes can be made without departing from theinvention.

What is claimed is:
 1. A biosensor apparatus for detecting a bindingevent between a ligand and ligand-binding receptor, comprisinganelectrode substrate having a detection surface, formed on said detectionsurface, a monolayer composed of hydrocarbon chains anchored at theirproximal ends to the detection surface; also anchored to the detectionsurface, a heterodimer-subunit complex composed of first and secondpeptides that together form an α-helical coiled-coil heterodimer, andsaid ligand, where(i) said first peptide is covalently bound to thedetection surface, at surface sites distinct from those at whichthe-hydrocarbon chains are anchored, (ii) said ligand is covalentlyattached to the second peptide, (iii) said hydrocarbon chains aresufficiently close-packed and ordered such that binding of the secondpeptide to the first peptide, to form said heterodimer-subunit complex,is effective to measurably reduce the electron flow across the monolayermediated by a redox ion species in an aqueous medium, relative toelectron flow observed in the presence of the first peptide alone, and(iv) binding of the ligand-binding receptor to the ligand is effectiveto measurably increase the electron flow across the monolayer mediatedby said redox ion species, a wall structure cooperating with saidsubstrate to define a chamber for holding the aqueous medium of theredox ion species in contact with said monolayer, means for introducingthe ligand-binding receptor into said chamber, and means for measuringsaid redox ion-mediated electron flow across said monolayer, in responseto said binding event occurring between said ligand-binding receptor andligand.
 2. The apparatus of claim 1, wherein the electrode comprises agold detection surface and said monolayer is composed of 8-22 carbonatom chains attached at their proximal ends to the detection surface bya thiol linkage.
 3. The apparatus of claim 1, wherein said chains have amolecular density of about 3 to 5 chains/nm².
 4. The apparatus of claim3, wherein the first peptide subunit is covalently anchored to thedetection surface through an oligopeptide spacer.
 5. The apparatus ofclaim 3, wherein the first peptide subunit is covalently anchored to thedetection surface through a hydrocarbon-chain spacer.
 6. The apparatusof claim 1, for use in detecting the presence, in a body-fluid sample,of a receptor which forms with said ligand, a ligand-receptor bindingpair selected from the group consisting of antigen-antibody,hormone-receptor, drug-receptor, cell-surface antigen-lectin,biotin-avidin, and complementary nucleic acid strands, wherein saidligand is selected from the group consisting of antigens, hormones,drugs, cell-surface antigens, and oligonucleotides.
 7. apparatus ofclaim 6, for use in detecting an antibody analyte, wherein the ligand isan antigen which forms an immunospecific binding pair with the antibodyanalyte.
 8. The apparatus of claim 6, for use in detecting asingle-stranded polynucleotide, wherein said ligand is anoligonucleotide whose sequence is complementary to a portion of thepolynucleotide sequence.
 9. The apparatus of claim 1, for use indetecting the presence, in a body-fluid sample, of an analyte whichforms with a receptor, an analyte-receptor binding pair wherein theapparatus further includes said receptor in said aqueous solution. 10.The apparatus of claim 9, wherein the analyte is an antigen, and thereceptor is an antibody effective to bind immunospecifically with saidanalyte and said ligand.